System and method for calculating a lumen pressure utilizing sensor calibration parameters

ABSTRACT

A system and method are provided for determining a pressure associated with a lumen of a body. A wireless sensor is positioned in the lumen of the body. The sensor comprises an LC resonant circuit having a resonant frequency configured to vary in response to changes in pressure in the lumen. One or more sensor calibration parameters are stored at an external base unit. The external based unit generates and transmits an energizing signal. A ring down response is received from the wireless sensor. The system and method determine the resonant frequency of the LC resonant circuit from the ring down response and calculate the pressure in the lumen from the resonant frequency of the LC resonant circuit utilizing the one or more sensor calibration parameters associated with the LC resonant circuit.

CROSS-REFERENCE TO RELATED APPLICATION

The present application is a divisional application of, and claimspriority to, U.S. application Ser. No. 17/401,365, filed 13-Aug.-2021entitled “Implantable Wireless Pressure Sensor and Method ofManufacture”. The present application is a divisional application of,and claims priority to, U.S. application Ser. No. 17/184,775, filed25-Feb.-2021 entitled “System and Method for Developing an ImplantAssembly”. The complete subject matter of each of the above identifiedapplications are expressly incorporated herein by reference in theirentirety.

The '365 application is a divisional application of, and claims priorityto, U.S. application Ser. No. 17/184,717, filed 25-Feb.-2021 entitled“Wireless Sensor for Measuring Pressure” (now U.S. Pat. No. 11,103,146).The '365 application is a divisional application of, and claims priorityto, U.S. application Ser. No. 17/184,755, filed 25-Feb.-2021 entitled“Method and System for Determining a Lumen Pressure” (now U.S. Pat. No.11,103,147). The 717, 755 and 775 applications are divisionalapplications of, and claims priority to, U.S. application Ser. No.16/194,103, filed 16-Nov.-2018 entitled “Wireless Sensor for MeasuringPressure” (now U.S. Pat. No. 11,033,192), which is a continuationapplication of U.S. application Ser. No. 14/733,450, filed 8-Jun.-2015entitled “Method of Manufacturing Implantable Wireless Sensor for InVivo Pressure Measurement” (now U.S. Pat. No. 10,143,388), which is acontinuation of U.S. application Ser. No. 12/612,070, filed 4-Nov.-2009entitled “Method of Manufacturing Implantable Wireless Sensor for InVivo Pressure Measurement” (now U.S. Pat. No. 9,078,563), which adivisional of U.S. application Ser. No. 11/204,812 filed 16-Aug.-2005entitled “Method of Manufacturing Implantable Wireless Sensor for InVivo Pressure Measurement” (now U.S. Pat. No. 7,621,036), which is acontinuation-in-part of U.S. application Ser. No. 11/157,375, filed21-Jun.-2005, entitled “Implantable Wireless Sensor for In Vivo PressureMeasurement” (now abandoned). The complete subject matter of each of theabove identified applications are expressly incorporated herein byreference in their entirety.

TECHNICAL FIELD

Embodiments herein relate to implantable sensors and methods ofmanufacturing implanted sensors for wirelessly sensing pressure,temperature and other physical properties within the human body. Moreparticularly, embodiments herein concerns a method of manufacturing awireless, un-powered, micromachined pressure sensor that can bedelivered using catheter-based endovascular or surgical techniques to alocation within an organ or vessel. Embodiments are further directed ingeneral to communicating with a wireless sensor, and in particular tocommunicating with a wireless sensor implanted within the body tomeasure a physical condition.

BACKGROUND

The measurement of blood pressure within the human heart and itsvasculature provides critical information regarding the organ'sfunction. Many methods and techniques have been developed to givephysicians the ability to monitor heart function to properly diagnoseand treat various diseases and medical conditions. For example, a sensorplaced within the chambers of the heart can be used to record variationsin blood pressure based on physical changes to a mechanical elementwithin the sensor. This information is then transferred through a wirefrom the sensor to an extracorporeal device that is capable oftranslating the data from the sensor into a measurable value that can bedisplayed. The drawback of this type of sensor is that there must be awired connection between the sensor and the extracorporeal device, thuslimiting Its use to acute settings.

Many types of wireless sensors have been proposed that would allowimplantation of the device into the body. Then, through the appropriatecoupling means, pressure reading can be made over longer periods ofinterest. The primary limitation to these type of sensors is that thefabrication methods used to manufacture them do not provide sufficientminiaturization to allow them to be introduced and implanted into theheart using nonsurgical, catheter based techniques while maintaining theability to communicate wirelessly with external electronics.

An implantable sensor of this type must be assembled using the materialsand fabrication methods that ensure appropriate biocompatibility andlong term mechanical and electrical durability.

One method of manufacturing a sensor capable of measuring pressure is touse a capacitor that is assembled such that one of the capacitive plateswill be displaced with respect to the other as a result of exposure toexternally applied stress. This displacement will result in a change inthe capacitance that is proportional to the applied stress. Variouspatents describe the fabrication and use of capacitor-based pressuresensors. The primary limitation of many of these inventions is that thetechniques used to fabricate the sensors do not lend themselves to theminiaturization necessary for it to be configured as an implantablemedical device while maintaining the capability of communicatingwirelessly with external electronics.

The fabrication methodologies that have been developed in the field ofMicro-Electro-Mechanical Systems (“MEMS”), however, do specificallyprovide the means for assembling miniaturized sensors capable ofmeasuring a variety of properties including pressure. MEMS devices asdescribed in prior patents traditionally use silicon as a substrate forconstruction of miniature electrical or mechanical structures.

A number of patents detail pressure sensors (some capacitive in nature,some manufactured using MEMS based fabrication methods) that arespecifically designed for implantation into the human body. Thesesensors suffer from many of the limitations already mentioned, with theadditional concerns that they require either the addition of a powersource to operate the device or the need for a physical connection to adevice capable of translating the sensor output into a meaningfuldisplay of a physiologic parameter.

To overcome the two problems of power and physical connection, theconcept of a externally modulated LC circuit has been applied todevelopment of implantable pressure sensors. Of a number of patents thatdescribe a sensor design of this nature, U.S. Pat. No, 6,113,553 toChubbuck is a representative example. The Chubbuck patent demonstrateshow a combination of a pressure sensitive capacitor placed in serieswith an inductor coil provides the basis for a wireless, un-poweredpressure sensor that is suitable for implantation into the human body.Construction of an LC circuit in which variations of resonant frequencycorrelate to changes in measured pressure and in which these variationscan be detected remotely through the use of electromagnetic coupling arefurther described in U.S. Pat. Nos. 6,111,520 and 6,278,379, both toAllen et al., incorporated herein by reference.

The device described in the Chubbuck patent is large, thus requiringsurgical implantation and thereby limiting its applicability to areasthat are easily accessible to surgery (e.g., the skull).

Thus, the need exists for a miniature, biocompatible, wireless,un-powered, hermetic pressure sensor that can be delivered into theheart or the vasculature using a small diameter catheter.

Further, U.S. Pat. Nos. 6,111,520, 6,855,115 and U.S. Publication No.2003/0136417, each of which is incorporated herein by reference, alldescribe wireless sensors that can be implanted within the body. Thesesensors can be used to monitor physical conditions within the heart oran abdominal aneurysm. An abdominal aortic aneurysm (AAA) is adilatation and weakening of the abdominal aorta that can lead to aorticrupture and sudden death. In the case of a repaired abdominal aneurysm,a sensor can be used to monitor pressure within the aneurysm sac todetermine whether the intervention is leaking. The standard treatmentfor AAAs employs the use of stent-grafts that are implanted viaendovascular techniques. However, a significant problem that has emergedwith these stent-grafts for AAAs is acute and late leaks of blood intothe aneurysms sac. Currently, following stent-graft implantation,patients are subjected to periodic evaluation via abdominal CT (ComputedTomography) with IV contrast to identify the potential presence ofstent-graft leaks. This is an expensive, risky procedure that lacksappropriate sensitivity to detect small leaks.

Typically, the sensors utilize an inductive-capacitive (“LC”) resonantcircuit with a variable capacitor. The capacitance of the circuit varieswith the pressure of the environment in which the sensor is located andthus, the resonant frequency of the circuit varies as the pressurevaries. Thus, the resonant frequency of the circuit can be used tocalculate pressure.

Ideally, the resonant frequency is determined using a non-invasiveprocedure. Several examples of procedures for determining the resonantfrequency of an implanted sensor are discussed in U.S. Pat. No.6,111,520. Some of the procedures described in the patent require thetransmission of a signal having multiple frequencies. A drawback ofusing a transmission signal having multiple frequencies is that theenergy in the frequency bands outside the resonant frequency is wasted.This excess energy requires more power which results in an increase incost, size, and thermal requirements, as well as an increase inelectromagnetic interference with other signals. Thus, there is a needfor an optimized method that is more energy efficient and requires lesspower.

There are unique requirements for communicating with an implantedsensor. For example, the system must operate in a low power environmentand must be capable of handling a signal from the sensor with certaincharacteristics. For example, the signal from the sensor is relativelyweak and must be detected quickly because the signal dissipates quickly.These requirements also impact the way that common problems are handledby the system. For example, the problems of switching transients andfalse locking need to be handled in a manner that accommodates thesensor signal characteristics. Thus, there is a need for a method forcommunicating with a wireless sensor that operates in a low powerenvironment and that efficiently determines the resonant frequency ofthe sensor.

The resonant frequency of the sensor is a measured parameter that iscorrelated with the physical parameter of interest. To be clinicallyuseful there must be means to ensure that variations in measurementenvironment do not affect the accuracy of the sensor. Thus, there is aneed for a system and method for communicating with a wireless sensorthat considers variations in the measurement environment.

SUMMARY

Stated generally, the present invention is directed toward a sensor andmethod for manufacturing a sensor to measure pressure within the heartor vasculature of a patient. The sensor comprises an upper wafer formedfrom a dielectric material, the upper wafer having one or more channels.The upper wafer includes a first capacitor plate and a second capacitorplate formed on a lower surface of the upper wafer. According to oneembodiment the sensor further comprises an inductor formed from one ormore windings of a conductive material, the inductor being containedwithin the one or more channels in the upper wafer in fixed relation tothe first and second capacitor plates, the inductor comprising first andsecond inductor leads, the first lead being electrically coupled to thefirst capacitor plate and the second lead electrically coupled to thesecond capacitor plate. The apparatus further comprises a lower waferformed from the dielectric material, the lower wafer being thinner thanthe upper wafer and a third capacitor plate formed on an inner surfaceof the lower wafer, the upper and lower wafers being fused together toform a monolithic housing such that the first and second capacitorplates are arranged in parallel, spaced-apart relation from the thirdcapacitor plate, a portion of the lower wafer comprising a pressuresensitive deflective region underlying at least a portion of the thirdcapacitor plate, whereby the deflective region deflects in response tochanges in ambient pressure in the medium.

Generally the invention further comprises a method for manufacturing asensor for measuring pressure within the heart or the vasculature of apatient by implanting a pressure sensor in such locations utilizingcatheter-based endovascular or surgical techniques and usingextracorporeal electronics to measure the pressure easily, safely, andaccurately. Stated somewhat more specifically, according to a firstaspect of manufacturing a sensor for in vivo applications, a recess isformed in a first wafer, and a capacitor plate is formed in the recessof the first wafer. A second capacitor plate is formed in acorresponding region of a second wafer. The two wafers are mutuallyimposed and affixed to one another such that the two capacitor platesare arranged in parallel, spaced-apart relation.

According to a second aspect of the invention, a method of manufacturinga sensor for in vivo applications comprises the step of providing threewafers of an electrically non-conductive material. First and secondcapacitor plates are formed on an upper surface of the first wafer. Athird capacitor plate is formed on a lower surface of the second wafer.The first and second wafers are then mutually imposed such that thethird capacitor plate is positioned in generally parallel, spaced-apartrelation from the first and second capacitor plates. An inductor coil ispositioned on top of an upper surface of the second wafer, and the leadsof the inductor coil are electrically connected to the first and secondcapacitor plates. A cavity is formed in the third wafer sufficient toreceive said inductor coil, and the third wafer is positioned on top ofthe second wafer with the inductor coil being received within the cavityof the third wafer. Finally, the second wafer is bonded to the first andthird wafers.

According to still another aspect of the invention, a method ofmanufacturing a sensor for in vivo applications, comprises the steps offorming a bottom plate on a wafer of electrically insulating material,forming a sacrificial layer over the bottom plate, forming a top plateon top of the sacrificial layer, and removing the sacrificial layer toleave the bottom and top plates in spaced-apart relation.

In yet another aspect of the present invention, a method ofmanufacturing a sensor for in vivo applications includes the step ofproviding first and second wafers. A recess is formed in the firstwafer, and a first plate is formed in the recess of the first wafer. Acoil-receiving trench is formed in an upper surface of the second wafer,and second and third plates are formed on the upper surface of thesecond wafer within the perimeter of the coil-receiving trench. Aninductor coil is positioned within the coil-receiving trench in theupper surface of the second wafer, and the leads of the inductor coilare electrically connected to the second and third plates on the uppersurface of the second wafer. The first and second wafers are affixed toone another such that the first plate in the recess of the first waferis in parallel, spaced apart relation to the second and third plates onthe upper surface of the second wafer.

Thus it is an object of this invention to provide a method formanufacturing an implantable wireless sensor.

It is also an object of this invention to provide a method formanufacturing a wireless, passive micromechanical sensor that can bedelivered endovascularly to a heart chamber or the vasculature.

It is a further object of this invention to provide a method formanufacturing an implantable, wireless, passive sensor that can bedelivered endovascularly to a heart chamber or the vasculature tomeasure pressure and/or temperature.

Other objects, features, and advantages of the present invention willbecome apparent upon reading the following specification, when taken inconjunction with the drawings and the appended claims.

Further, a goal of aneurysm treatment is to depressurize the sac and toprevent rupture. Endoleaks, whether occurring intraoperatively orpostoperatively, can allow the aneurysmal sac to remain pressurized andtherefore, increase the chance of aneurysm rupture. The current imagingmodalities angiography and CT scan are not always sensitive enough todetect endoleaks or stent graft failure. Intrasac pressure measurementsprovide a direct assessment of sac exclusion from circulation and maytherefore offer intraoperative and post-operative surveillanceadvantages that indirect imaging studies do not.

In applications of embodiments herein, an AAA pressure sensor is placedinto the aneurysm sac at the time of stent-graft insertion. The pressurereadings are read out by the physician by holding an electronicinstrument, which allows an immediate assessment of the success of thestent-graft at time of the procedure and outpatient follow-up visits, byreading the resonant frequency of the wireless sensor and correlatingthe frequency reading to pressure.

Embodiments herein meets the needs described above by providing a systemand method for communicating with a wireless sensor to determine theresonant frequency of the sensor. The system energizes the sensor with alow duty cycle, gated burst of RF energy having a predeterminedfrequency or set of frequencies and a predetermined amplitude. Theenergizing signal is coupled to the sensor via a magnetic loop. Thesensor may be an inductive-capacitive (“LC”) resonant circuit with avariable capacitor that is implanted within the body and used to measurephysical parameters, such as pressure or temperature. The energizingsignal induces a current in the sensor which is maximized when theenergizing frequency is the same as the resonant frequency of thesensor. The system receives the ring down response of the sensor viamagnetic coupling and determines the resonant frequency of the sensor,which is used to calculate the measured physical parameter.

In one aspect, a pair of phase locked loops (“PLLs”) is used to adjustthe phase and the frequency of the energizing signal until its frequencylocks to the resonant frequency of the sensor. In one embodiment, onePLL samples during the calibration cycle and the other PLL samplesduring the measurement cycle. These cycles alternate every 10microseconds synchronized with the pulse repetition period. Thecalibration cycle adjusts the phase of the energizing signal to a fixedreference phase to compensate for system delay or varying environmentalconditions. The environmental conditions that can affect the accuracy ofthe sensor reading include, but are not limited to, proximity ofreflecting or magnetically absorbpative objects, variation of reflectingobjects located within transmission distance, variation of temperatureor humidity which can change parameters of internal components, andaging of internal components.

One of the PLLs is used to adjust the phase of the energizing signal andis referred to herein as the fast PLL. The other PLL is used to adjustthe frequency of the energizing signal and is referred to herein as theslow PLL. During the time that the energizing signal is active, aportion of the signal enters the receiver and is referred to herein as acalibration signal. The calibration signal is processed and sampled todetermine the phase difference between its phase and the phase of alocal oscillator (referred to herein as the local oscillator 2). Thecycle in which the calibration signal is sampled is referred to as thecalibration cycle. The system adjusts the phase of the energizing signalto drive the phase difference to zero or another reference phase.

During the measurement cycle, the signal coupled from the sensor(referred to herein as the coupled signal or the sensor signal) isprocessed and sampled to determine the phase difference between thecoupled signal and the energizing signal. The system then adjusts thefrequency of the energizing signal to drive the phase difference to zeroor other reference phase. Once the slow PLL is locked, the frequency ofthe energizing signal is deemed to match the resonant frequency of thesensor. The operation of the slow PLL is qualified based on signalstrength so that the slow PLL does not lock unless the strength of thecoupled signal meets a predetermined signal strength threshold.

The system also handles false locking and switching transients. A falselock occurs if the system locks on a frequency that does not correspondto the resonant frequency of the sensor. In one aspect of the invention,the system avoids false locks by examining how the phase differencesignal goes to zero. If the slope of the phase difference signalrelative to time meets a predetermined direction, e.g. positive, thenthe PLL is allowed to lock. However, if the slope of the phasedifference signal relative to time does not meet the predetermineddirection, e.g. it is negative, then the signal strength is suppressedto prevent a false lock.

Another aspect herein uses frequency dithering to avoid a false lock. Aconstant pulse repetition frequency can add spectral components to thesensor signal and cause a false lock. By randomly varying the pulserepetition frequency of the energizing signal, the sidebands move backand forth so that the average of the sidebands is reduced. Thus, thesystem locks on the center frequency rather than the sidebands.

In another aspect, phase dithering can be used to reduce switchingtransients. The phase of the energizing signal and a local oscillator(referred to herein as local oscillator 1) are randomly changed. Varyingthe phase of the energizing signal varies the phase of the coupledsignal, but does not affect the phase of the transient signal. Thus, theaverage of the transient signal is reduced. Changing the resonantfrequency of the coil as it is switched from energizing mode to couplingmode also reduces switching transients. The capacitors that areconnected to the coil are switched between different modes to slightlychange the resonant frequency in order to reduce switching transients.

These and other aspects, features and advantages may be more clearlyunderstood and appreciated from a review of the following detaileddescription of the disclosed embodiments and by reference to theappended drawings and claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of a first embodiment of an implantablewireless sensor according to the present invention, with the sensor bodyshown as transparent to reveal interior detail.

FIG. 2 is a schematic view of two pressure sensitive capacitor platesbeing formed in recessed trenches on two substrate wafers.

FIG. 3 is a schematic view showing the wafers of FIG. 2 imposed inface-to-face relation.

FIG. 4 is a schematic view showing the imposed wafers of FIG. 3 beinglaser-cut around their peripheries.

FIG. 5 is a schematic view of an alternate embodiment of two imposedwafers in which only one of the wafers has a recessed trench.

FIG. 6 is a schematic view illustrating a first step in a process formanufacturing wafers with capacitor plates formed thereon.

FIG. 7 is a schematic view illustrating a second step in a process formanufacturing wafers with capacitor plates formed thereon.

FIG. 8 is schematic view illustrating a third step in a process ormanufacturing wafers with capacitor plates formed thereon.

FIG. 9 is a schematic view illustrating a fourth step in a process formanufacturing wafers with capacitor plates formed thereon.

FIG. 10 shows another embodiment in which two capacitor plates areformed on one wafer.

FIG. 11 Illustrates the embodiment of FIG. 10 showing the two capacitorplates on the single wafer connected to opposite ends of a helicalinductor coil.

FIG. 12 is a schematic view of still another embodiment of animplantable, wireless pressure sensor.

FIG. 13 is a schematic view of a further embodiment of an implantable,wireless pressure sensor in which a three-dimensional inductor coil isbuilt onto the top of through connection terminals on the backside of acapacitor plate substrate.

FIG. 14 is a schematic view of another embodiment of a wireless pressuresensor in which each subsequent layer is alternately spaced slightlysmaller or larger in diameter than the previous winding.

FIG. 15 is a schematic view of a further embodiment of an implantable,wireless pressure sensor in which a three-dimensional inductor coil isbuilt onto the surface of a cylinder.

FIG. 16 is a schematic view of another embodiment of a wireless pressuresensor in which the pressure sensitive capacitor and three-dimensionalinductor coil are formed together on one wafer.

FIG. 17 is a schematic view showing a first step in the manufacturingprocess of the wireless pressure sensor of FIG. 16.

FIG. 18 is a schematic view showing a second step in the manufacturingprocess of the wireless pressure sensor of FIG. 16.

FIG. 19 is a schematic view showing a third step in the in amanufacturing process of the wireless pressure sensor of FIG. 16.

FIG. 20 is a schematic view showing a fourth step in the manufacturingprocess of the wireless pressure sensor of FIG. 16.

FIG. 21 is a schematic view showing a fifth step in the manufacturingprocess of the wireless pressure sensor of FIG. 16.

FIG. 22 shows a first arrangement for electrically and mechanicallyinterconnecting a capacitor plate to an inductor coil.

FIG. 23 shows a second arrangement for electrically and mechanicallyinterconnecting a capacitor plate to an inductor coil.

FIG. 24 is a schematic view of another embodiment of a wireless pressuresensor in which the pressure sensitive capacitor and three-dimensionalinductor coil are formed on two wafers.

FIG. 25 is a schematic view showing a first step in the manufacturingprocess of the wireless pressure sensor of FIG. 24.

FIG. 26 is a schematic view showing a second step in the manufacturingprocess of the wireless pressure sensor of FIG. 24.

FIG. 27 is a schematic view showing a third step in the manufacturingprocess of the wireless pressure sensor of FIG. 24.

FIG. 28 is a schematic view showing a fourth step in the manufacturingprocess of the wireless pressure sensor of FIG. 24.

FIG. 29 is a schematic view of an embodiment of a wireless pressuresensor utilizing four wafers.

FIG. 30 is a schematic view showing a first step in the manufacturingprocess of the wireless pressure sensor of FIG. 29.

FIG. 31 is a schematic view showing a second step in the manufacturingprocess of the wireless pressure sensor of FIG. 29.

FIG. 32 is a schematic view showing a third step in the manufacturingprocess of the wireless pressure sensor of FIG. 29.

FIG. 33 is a side view of a pressure sensor and a retention mechanism ofa delivery device, with the retention mechanism in a dosedconfiguration.

FIG. 34 is a side view of the pressure sensor and retention mechanismFIG. 33, with the retention mechanism in an open configuration.

FIG. 35 is a side view of the pressure sensor and retention mechanismFIG. 33, with the retention mechanism in a dosed configuration and shownin cross-section.

FIG. 36 is a side view of the pressure sensor and retention mechanismFIG. 33, with the retention mechanism in an open configuration and shownin cross-section.

FIG. 37 is a side view of a dual-coil shaft of a delivery device, with aportion of the outer coil being removed to show the inner coil.

FIG. 38 is a side view of a delivery device comprising the retentionmechanism of FIG. 33 and the shaft of FIG. 37, illustrating a first stepin the delivery of a sensor into the wall of a septum.

FIG. 39 is a side view of the delivery device of FIG. 38, illustrating asecond step in the delivery of a sensor into the wall of a septum.

FIG. 40 is a side view of the delivery device of FIG. 38, illustrating athird step in the delivery of a sensor into the wall of a septum.

FIG. 41 is a side view of the delivery device of FIG. 38, illustrating afourth step in the delivery of a sensor into the wall of a septum.

FIG. 42 is a side view of an alternate embodiment of a delivery devicefor delivering a sensor into the wall of a septum, with the retentionmechanism of the delivery device in a closed configuration.

FIG. 43 is a side view of the delivery device of FIG. 42 showing theretention mechanism in an open configuration.

FIG. 44 is an isometric view of a sensor comprising an alternatearrangement for anchoring the sensor within a lumen of a patient.

FIG. 45 is a top view of the sensor of FIG. 44.

FIG. 46 is a top view showing the sensor of FIG. 44 lodged within alumen.

FIG. 47 is a side cutaway view of a shaft of a delivery apparatus forimplanting the sensor of FIG. 44.

FIG. 48 is a side view of a tether wire of a delivery apparatus forimplanting the sensor of FIG. 44.

FIG. 49 is a side view of a core wire of a delivery apparatus forimplanting the sensor of FIG. 44.

FIG. 50 is a side view of a guidewire of a delivery apparatus forimplanting the sensor of FIG. 44.

FIG. 51 is a side cutaway view of a delivery apparatus comprising thecomponents of FIGS. 47-50 with the sensor of FIG. 44 mounted thereto.

FIG. 52 is a block diagram of an exemplary system for communicating witha wireless sensor in accordance with an embodiment of the invention.

FIG. 53A is a graph illustrating an exemplary energizing signal inaccordance with an embodiment of the invention.

FIGS. 53B, 53C and 53D are graphs illustrating exemplary coupled signalsin accordance with an embodiment of the invention.

FIG. 54 is a block diagram of an exemplary base unit in accordance withan embodiment of the invention.

FIGS. 55A and 55B are graphs illustrating exemplary phase differencesignals in accordance with an embodiment of the invention.

FIG. 56 illustrates frequency dithering in accordance with an embodimentof the invention.

FIG. 57 illustrates phase dithering in accordance with an embodiment ofthe invention.

FIG. 58 illustrates a coupling loop in accordance with an embodiment ofthe invention.

FIG. 59 is a graph illustrating an exemplary charging response of an LCcircuit in accordance with an embodiment of the invention.

DETAILED DESCRIPTION

Referring now to the drawings, in which like numerals indicate likeelements throughout the several views, FIG. 1 illustrates a sensor 10for the measurement of physical parameters. The sensor can be fabricatedusing micro-machining techniques and is small, accurate, precise,durable, robust, biocompatible, and insensitive to changes in bodychemistry, or biology. Additionally, the sensor can incorporateradiopaque features to enable fluoroscopic visualization duringplacement within the body. Furthermore, this sensor is encased in ahermetic, unitary package of electrically insulating material where thepackage is thinned in one region so as to deform under a physiologicallyrelevant range of pressure. The LC circuit contained in the packaging isconfigured so that one electrode of the capacitor is formed on thethinned region. This sensor does not require the use of externalconnections to relay pressure information externally and does not needan internal power supply to perform its function. The pressure sensor ofthe current invention can be attached to the end of a catheter to beintroduced into a human body and delivered to an organ or vessel usingcatheter-based endovascular techniques.

Referring to FIG. 1, the sensor 10 includes a body 12. The body 12 isformed from electrically insulating materials, preferably biocompatibleceramics. In a preferred embodiment, the body is comprised of fusedsilica. The sensor 10 comprises a deflectable region 14 at the lower endof the body 12. The body 12 further comprises a lower chamber 19 and anupper chamber 21.

An LC resonator is hermetically housed within the body 12 and comprisesa capacitor 16 and an inductor 20. As used herein, the term “hermetic”will be understood to mean “completely sealed, especially against theescape or entry of air and bodily fluids. ” The capacitor 15 is locatedwithin the lower cylindrical chamber 19 and comprises at least twoplates 16, 18 disposed in parallel, spaced apart relation. The inductor20 comprises a coil disposed within the upper chamber 21 and which is inconductive electrical contact with the capacitor 15.

The lower capacitor plate 18 is positioned on the inner surface of thedeflectable region 14 of the sensor body 12. The upper capacitor plate16 is positioned on a fixed region of the sensor body 12. A change inambient pressure at the deflectable region 14 of the sensor 10 causesthe deflectable region 14 to bend, thereby displacing the lower plate 16with respect to the upper plate 18 and changing the capacitance of theLC circuit. Because the change in capacitance of the LC circuit changesits resonant frequency, the resonant frequency of the sensor 10 ispressure-dependent.

Beyond what has been presented in U.S. Pat. Nos. 6,111,520 and 6,278379,covering the fundamental operating principle of the wireless pressuresensor, additional means to further sensor miniaturization is requiredin order to achieve an acceptable size for implantation into the heartor the vasculature. The sensor outer dimensions are constrained by thelumen size of the delivery catheter that is used to introduce thesensor. Catheter inner diameters typically range from 1-5 mm. Also, thesize and shape of the sensor should minimally interfere with mechanicalor hemodynamic function of the heart or vessel where it is located.

Within these physical size constraints, one of the most significantchallenges is achieving adequate coupling to the sensor inductor coilfrom the external readout device at the necessary distance from theoutside of the body to the implant site. One method for achievingenhanced coupling is to add magnetic material to the inductor. However,this approach is not feasible in a sensor intended for in vivo use, asthe magnetic material would be adverse to magnetic resonance imaging,for example. For a limited coil cross-sectional area, an increasedcoupling coefficient is also achievable by using a three-dimensionalinductor coil configuration, as opposed to two-dimensional designs. Forthese reasons, a three-dimensional helical inductor coil configuration20 is the preferred embodiment for the sensor design.

The disclosed sensor features a completely passive inductive-capacitive(LC) resonant circuit with a pressure varying capacitor. Because thesensor is fabricated using completely passive electrical components andhas no active circuitry, it does not require on-board power sources suchas batteries, nor does it require leads to connect to external circuitryor power sources. These features create a sensor which is self-containedwithin the packaging material and lacks physical interconnectionstraversing the hermetic packaging, such interconnects frequently beingcited for failure of hermeticity. Furthermore, other sensingcapabilities, such as temperature sensing, can be added using the samemanufacturing techniques. For example, temperature sensing capabilitycan be accomplished by the addition of a resistor with known temperaturecharacteristics to the basic LC circuit.

The capacitor in the pressure sensor of the disclosed invention consistsof at least two conductive elements separated by a gap. If a force isexerted on the sensor, a portion of the sensor deflects, changing therelative position between the at least two conductive elements. Thismovement will have the effect of reducing the gap between the conductiveelements, which will consequently change the capacitance of the LCcircuit. An LC circuit is a closed loop system whose resonance isproportional to the inverse square root of the product of the inductorand capacitor. Thus, changes in pressure alter the capacitance and,ultimately, cause a shift in the resonant frequency of the sensor. Thepressure of the environment external to the sensor is then determined byreferencing the value obtained for the resonant frequency to apreviously generated curve relating resonant frequency to pressure.

Because of the presence of the inductor, it is possible to couple to thesensor electromagnetically and to induce a current in the LC circuit viaa magnetic loop. This characteristic allows for wireless exchange ofelectromagnetic energy with the sensor and the ability to operate itwithout the need for an on-board energy source such as a battery. Thusit is possible to determine the pressure surrounding the sensor by asimple, non-invasive procedure by remotely interrogating the sensor,recording the resonant frequency, and converting this value to apressure measurement.

One method of sensor interrogation is explained in U.S. patentapplication Ser. No. 11/105,294, incorporated herein by reference.According to this invention, the interrogating system energizes thesensor with a low duty cycle, gated burst of RF energy having apredetermined frequency or set of frequencies and a predeterminedamplitude. The energizing signal is coupled to the sensor via a magneticloop. The energizing signal induces a current in the sensor that ismaximized when the frequency of the energizing signal is substantiallythe same as the resonant frequency of the sensor. The system receivesthe ring down response of the sensor via magnetic coupling anddetermines the resonant frequency of the sensor, which is then used todetermine the measured physical parameter. The resonant frequency of thesensor is determined by adjusting the frequency of the energizing signaluntil the phase of the ring down signal and the phase of a referencesignal are equal or at a constant offset. In this manner, the energizingsignal frequency is locked to the sensors resonant frequency and theresonant frequency of the sensor is known. The pressure of the localizedenvironment can then be ascertained.

Q factor (Q) is the ratio of energy stored versus energy dissipated. Thereason Q is important is that the ring down rate of the sensor isdirectly related to the Q. If the Q is too small, the ring down rateoccurs over a substantially shorter time interval. This necessitatesfaster sampling intervals, making sensor detection more difficult. Also,as the Q of the sensor increases, so does the amount of energy returnedto external electronics. Thus, it is important to design sensors withvalues of Q sufficiently high enough to avoid unnecessary increases incomplexity in communicating with the sensor via external electronics.

The Q of the sensor is dependent on multiple factors such as the shape,size, diameter, number of turns, spacing between the turns andcross-sectional area of the inductor component. In addition Q will beaffected by the materials used to construct the sensors. Specifically,materials with low loss tangents will provide a sensor with higher Qfactors.

The body of the implantable sensor of the disclosed embodiment of thepresent invention is preferably constructed of ceramics such as, but notlimited to, fused silica, quartz, pyrex and sintered zirconia, thatprovide the required biocompatibility, hermeticity and processingcapabilities. These materials are considered dielectrics, that is, theyare poor conductors of electricity but are efficient supporters ofelectrostatic or electroquasistatic fields. An important property ofdielectric materials is their ability to support such fields whiledissipating minimal energy. The lower the dielectric loss, the lower theproportion of energy lost, and the more effective the dielectricmaterial is in maintaining high Q.

With regard to operation within the human body, there is a secondimportant issue related to Q, namely that blood and body fluids areconductive mediums and are thus particularly lossy. As a consequence,when a sensor is immersed in a conductive fluid, energy from the sensorwill dissipate, substantially lowering the Q and reducing thesensor-to-electronics distance. It has been found that such loss can beminimized by further separation of the sensor from the conductiveliquid. This can be accomplished, for example, by coating the sensor ina suitable low-loss-tangent dielectric material. The potential coatingmaterial must also meet stringent biocompatibility requirements and besufficiently compliant to allow transmission of fluid pressure to thepressure-sensitive deflective region. One preferred material for thisapplication is silicone rubber. It should be appreciated that use of acoating is an optional feature and is not required to practice theinvention per se but such coatings will preserve the Q of the sensorwhich can prove advantageous depending on the intracorporeal location ofthe sensor.

There are various manufacturing techniques that can be employed torealize sensors according to the current invention. Capacitors andinductors made by a variety of methods can be manufactured separately,joined through interconnect methods and encapsulated in hermeticpackaging. In one embodiment, the pressure sensitive capacitor 15 andthe three-dimensional inductor coil 20 are formed separately and joinedtogether to form the LC circuit. In another embodiment, the capacitorand inductor coil can be manufactured integral with one another.Additionally, there are several methods to create these discreteelements and to join each discrete element to create the final sensor.The following examples are provided to illustrate important designconsiderations and alternative methods for creating these discretesensor elements but should not be construed as limiting the invention inany way.

Referring to FIG. 12, the inductor coil 320 is comprised of the inductorcoil body 322 and the coli leads 324. Numerous parameters of theinductor coil can be varied to optimize the balance of size andelectrical properties of the circuit, including the materials, coildiameter, wire gage, insulation thickness, number of coil windings, andcross-sectional area of the coil body. The material comprising the coilmust be highly conductive and also biocompatible. Suitable materialsinclude, but are not limited to, gold, copper, and alloys thereof.

It is preferable in the practice of the disclosed invention to minimizeor eliminate changes in resonant frequency of sensors of the inventiondue to factors other than capacitance in order to reliably correlate theshift in resonant frequency with a change in distance between thecapacitor plates. Thus, it is important that the inductor coil 320 insensors of the current invention maintain a high degree of mechanicalstability as a change in coil position relative to the capacitor or achange in coil configuration will cause the resonant frequency of thedevice to change. There are many ways to immobilize the inductor coil320 of the present invention. If the wire used to construct the coil issufficiently strong, the coil can be self-supporting, also known as an“air core” configuration. A solenoid coil is another suitableconfiguration. If the wire is not sufficiently strong to maintain itsintended configuration during assembly and in use, the coil can beformed around a central bobbin comprised of a suitable material. Suchbobbins can be configured to be mechanically fixed to any surface orcombination of surfaces defining the coil receiving trench via a pressfit. Alternatively, the coil can be wound on a thermoplastic bobbinwhere the thermoplastic material can be subjected to sufficient heat tocause flow to encapsulate and/or adhere to the surface of the coilreceiving trench.

Alternatively, a thermosetting or thermoplastic polymer with good hightemperature characteristics, low loss tangent, and, optionally, lowdielectric constant material can be used to support the coil. Thepolymer should also be highly inert, have excellent aging resistance andexhibit substantially no moisture absorbance or outgassing. With the useof a thermosetting material, the polymer is applied to the coil inliquid form and allowed to cure or otherwise harden. Thermoplasticmaterials can be preformed and inserted between the coil and at leastone coil receiving trench wall and subsequently heated to achievesufficient flow to encapsulate and/or adhere to the coil and at leastone coil receiving trench wall.

Polyimide, fluorinated polymers, glass frit, ceramic paste and liquidcrystal polymer are examples of suitable materials for immobilizing theinductor coil 320 due to their thermal, electrical, and mechanicalproperties. However, manufacturing processes achieving substantiallysimilar results that involve lower processing temperatures would makeother material choices desirable, such choices being obvious to oneskilled in the art.

The wire from which the coil is formed can be solid wire, bundled wireor cable, or individually insulated stranded wire.

The wire gage, coil diameter, cross-sectional area of the coil body, andnumber of windings all influence the value of inductance and thedetection range of the circuit. As any of these properties increase, sodo the size and the inductance of the coil, as well as thesensor-to-electronics distance. To specify an inductor coil for use inthe sensor, size considerations must be balanced with those ofinductance and Q.

A small scale three-dimensional inductor coil can be formed in a varietyof ways. It can be created conventionally. One such method is machinecoil winding of small diameter insulated magnet wire, as shown in FIG.1.

In another embodiment, shown in FIG. 13, a three-dimensional inductorcoil 420 is built onto the top of one of the through connectionsterminals 480 on the backside of the capacitor plate substrate 442,using integrated circuit processing techniques and a multitude oflayers. This coil 420 can be defined and supported by photo-definabledielectric material such as photo-definable polyimide. In the disclosedembodiment, the coil is free standing in air, supported by same-materialmechanical elements that are strategically positioned to minimize theeffect of the supporting mechanical elements on the electrical functionof the coil.

In this approach it is desirable to minimize the number of design layersto improve batch process yield and to reduce processing time. In aconventional configuration, such as that shown in FIG. 13, a spacinglayer is required between each winding, making the number of layersrequired equal to two times the number of windings. In one version 500of the three-dimensional coil design, an example of which is shown inFIG. 14, each subsequent coil 510 is alternately spaced slightly smalleror larger in diameter than the previous winding. This configurationcreates a small separation between adjacent coils 510 in the x-y plane,eliminating the need for an extra vertical spacing layer in betweenwindings. This configuration results in a number of coil windings equalto the number of layers, which is more practical for manufacturing usinga MEMS approach.

In yet another embodiment 550, shown in FIG. 15, a three-dimensionalinductor coil 555 is built onto the surface of a cylinder 560 of anappropriate material such as, but not limited to fused silica. Aconductive layer is first applied to the surface of the cylinder 560.Then a mold is formed onto the surface so that parts of the underlyingconductive surface are exposed and some are covered. A metal may then beformed onto the exposed areas by electroplating, sputtering or vapordeposition. The exposed area forms a helical trench that extends alongthe surface of the cylinder, thus realizing an inductor coil.

Referring now to FIG. 2, the pressure sensitive capacitor plates 16, 18are formed on two separate substrate wafers 40, 42 in recessed trenches44. At least one of the wafers 40 has a substrate thickness in theregion 46 of the capacitive plate 16 such that sufficient patedeflection occurs due to external pressure change, resulting in asufficient change in resonant frequency per unit pressure (mm Hg) oncethe LC circuit has been created. If necessary, the thickness of thewafer 40 in the region 46 can be reduced by suitable chemical ormechanical means, as indicated by the dashed line 47, to provide thedesired range of deflection.

As shown in FIG. 3, the wafers 40, 42 are bonded together such that thecapacitive plates are 16, 18 parallel and separated by a gap on theorder of 0.1-10 microns, preferably 0.1-2 microns.

The performances of the sensor, especially the propensity of itscapacitance and, in turn, its resonant frequency to change as a responseto an environmental pressure change, are closely related to fewfundamental geometrical considerations. Widening or elongating thedeflective region will augment its mechanical flexibility, and, in turn,the pressure sensitivity of the sensor. Decreasing the thickness of thedeflective area will result in similar improvements. However, thinnerdeflective region can become too fragile or otherwise more sensitive tosystemic response from the host-organism other than changes in mean andpulsatile blood pressure (ex: hyperplasia, tissue overgrowth, etc.).Reducing the gap, while maintaining adequate deflective regionthickness, offers a complementary alternative to insufficiently lowsensitivity. As the initial value of the gap is shrinking, the motion ofthe deflective region relative to the initial gap becomes proportionallymore important. This results in a greater change in capacitance for agiven stimulus, therefore enhancing the pressure sensitivity. Whilerelevant sensitivity can be achieved with initial air-gap ranging from0.1 to 10 micrometers, initial air-gaps ranging from a 0.1 to 2micrometers are preferable.

To ensure adequate pressure range, the value of the maximum deflectionunder maximum load (indexed, for example, on physiologically relevantmaximum pulsatile blood pressure values, at relevant location in thehost-organism) ought to be, in theory, inferior or equal to the value ofthe initial gap. In practice, limiting the maximum deflection undermaximum bad to represent only a fraction of the initial gap (ex: 0.6micrometer for a 1 micrometer initial gap) will ease the fabricationconstraints and result in a more robust and versatile sensor.

One suitable method for creating the pressure sensitive capacitor is byelectroplating the individual plates 16, 18 in the recessed trenches 44on a substrate wafer 40, 42 to a given height H1, H2 that is less thanor equal to the depth D1, D2 of the respective trench 44. When thewafers are bonded together the capacitive plates are generally separatedby the difference between the sum of the trench depths and the sum ofthe plate heights, (D1+D2)-(H1+H2). An inherent variation in the heightof the plates and the required range of deflection for the fulloperating pressure range are parameters, which determine the initialseparation distance (a.k.a. the gap).

FIG. 4 illustrates the assembled wafers and capacitor plates laser-cutaround their peripheries 48, reducing the capacitor to its final sizeand hermetically fusing the two wafers together at 50. A CO.sub.2 lasercan be used at a peak wavelength of about 10 microns if the substrate isfused silica. Power must be sufficiently large to cut and fuse thewafers together, while at the same time being sufficiently small thatthe internal components of the sensor are not damaged by excessive heat.

In an alternate method, the wafers are pre-bonded using glass frit toproduce a hermetic seal around the cavities. In this method, the lasercut only releases the sensors from the wafer, and does not provide theprimary means of creating the hermetic seal. Other suitable methods ofhermetically sealing the wafers include, but are not limited to,adhesives, gold compression bonding, direct laser bonding, and anodicbonding.

In an alternate embodiment illustrated in FIG. 5, one plate 18 is formedon a substrate wafer 142 having a trench 144 with a depth greater thatof the trench 44 in the substrate wafer 40. The other plate 16 is formedon the inner surface of a wafer 140 without a trench. When imposed inface-to-face relation, the plate 16 is received into the lower end ofthe trench 144 with the plates 16, 18 disposed in parallel, spaced-apartrelation.

To achieve smaller gap separation distances on the order of 0.1-2microns, revised processing methods are employed to bring additionalcontrol to the variation in height across the conductive plates 16, 18.One method is as follows: the conductive plate 16, 18 is built to atarget height that slightly exceeds the depth of the recess trench 44,as shown in FIG. 6. In the disclosed embodiment the plates are formed byelectroplating. Preferred materials for the plates are copper, gold, andalloys thereof. After building the plates, each conductive plate 16, 18is polished using chemical/mechanical polishing (CMP) to planarize andreduce the height of the plate until it is less than the depth of thetrench by the desired amount, as shown in FIG. 9.

Another method also begins with the plates 16, 18 formed to a heightthat slightly exceeds the depth of the trenches 44, as shown in FIG. 6.The metal capacitor plates 16, 18 are mechanically polished to planarizethe metal surface down to the surface of the substrate 40, 42, as shownin FIG. 7. Following this step, the metal plates are chemically etchedby a selective etchant to the height indicated by the dashed line 56 inFIG. 8 to achieve the desired difference in height between the height ofthe plate 16, 18 and the depth of the trench 44, as shown in FIG. 9.

Still another method for forming the plates is physical vapor deposition(PVD), also known as thin film deposition, in conjunction withphotolithography. PVD is used to deposit a uniform layer of metal,sub-micrometer to tens of micrometers thick, on a wafer. Subsequently alayer of photoresist is deposited, a mask is used to pattern thephotoresist, and a selective etching technique is utilized to etch awaythe extra metal and to define the desired pattern. Other methods ofdefining the metal pattern can be utilized, such as, shadow masking, amethod well known in the art.

In one approach, shown in FIGS. 10 and 11, a pressure sensitivecapacitor 215 can be formed by separating the bottom conductive pad intotwo separate regions 218A, 2186 that capacitively couple to one anothervia a common third conductive region 216 on the pressure sensitivedeflective region. The inductor coil 20 is then electrically connectedas shown in Ha 11, one lead 22 of the coil 20 to the first region 218A,and the other lead 24 of the coil 20 to the second region 2186.

When the split-plate design is employed for one side of the capacitor,as shown in FIG. 11, the spat plates 218A, 218B are preferably locatedon the fixed side of the capacitor (i.e., opposite thepressure-sensitive side), because the electrical/mechanicalinterconnects made to the spot plates in order to complete the LCcircuit are less prone to mechanical failure when the surface to whichthey are mechanically attached does not deflect or move repetitively.

In yet another embodiment, shown in FIG. 12, the plate on the top wafer42 is separated by a dielectric into two conductive regions 318A, 318B,with one region 318B substantially larger than the other 318A. Afterbonding together of the two wafers 40, 42, the smaller conductive region318A is electrically connected to the outer edge of the pressuresensitive plate 316, spanning the air gap with a laser weld that isperformed through the substrate material. The laser wavelength isselected so that it is passes through the substrate material withminimal energy absorption, but heats the conductive plate sufficientlyto produce the weld connection between the top and bottom plates 316,318A.

It will be appreciated that sensors embodied by the current inventioncan have capacitive and inductive elements maintained in separatehermetic cavities or that these elements may be contained in a singlehermetic cavity.

In one embodiment, the pressure sensitive capacitor 15 needs to beconnected to the three-dimensional inductor coil 20 while maintaining ahermetic seal around the internal cavity that defines the separation gapbetween the capacitive plates 16, 18. This can be achieved by using avariety of through-wafer interconnection methods, familiar to thoseskilled in the art. Referring to FIG. 22, through holes or vias 660 areformed in an upper wafer 662 to provide mechanical and electrical accessto a pair of upper capacitor plates 664, 666. The wafer through-holescan be formed before or after plate formation using some combination ofthe following techniques: laser drilling, chemical (wet) etching,conventional or ultrasonic machining, or dry etching. As shown in FIG.22, the vies 660 can optionally be filled with gold, copper, or othersuitable conductive material to form through-wafer interconnects 668 inconductive communication with the capacitor plates 664, 666. Thethrough-wafer interconnects 668 thus form a hermetic seal. Leads from aninductor coil (not shown) are attached to the through-waferinterconnects 668 to place the leads m conductive communication with thecapacitor plates 664, 666.

Referring to FIG. 23, through holes or vies 680 are formed In an upperwafer 682 to provide mechanical and electrical access to a pair of lowercapacitor plates 684, 686. Electrical connections to the lower capacitorplates 684, 686 will be accomplished through leads of the inductor coil(not shown) or through wires or other suitable conductive means.

Thermosonic or ultrasonic bonding can be used to connect the inductorcoil to either an electrode of a capacitor or a through-waferinterconnect. Thermosonic and ultrasonic bonding are types of wirebonding used for metal wires including, but not limited to, gold wires.Typical temperatures required for thermosonic bonding are between125-220.degree. C., and bonding occurs when a combination of static andultrasonic mechanical and thermal energy is delivered to the metalliccoil wire to be bonded to a metal surface. Ultrasonic bonding isperformed just as thermosonic bonding but without the use of heat.Useful materials for the metallized bond sites and coil comprise gold,copper and aluminum and alloys thereof. Bonds can be formed betweencertain dissimilar metals as well as between all like metals, and suchcombinations are widely known in the art.

If the metal or metal alloy used for the coil has a dielectric (e.g.,polymer) coating, the coating must be removed prior to bonding. Thecoating can be removed to expose the metal at the adhesion point so thatbonding can occur by either mechanical or chemical means. Alternatively,the parameters (e.g. time, heat, pressure) of the thermosonic bondingprocess can be altered and the geometry of the bonding tool modified sothat reliable mechanical and electrical interconnects are created. Suchmodifications cause the coating material to be pushed aside, exposingthe metal at the bonding site and extruding the wire slightly. Thislatter technique provides certain advantages because it reduces thenumber of manufacturing steps.

An alternate method of conductively connecting the coil to thecapacitive plates is the solder bump. Solder is applied to themetal-metal interface of the coil and electrode or interconnect to forma mechanical and electrical connection. This method can be used forcapacitor plate or through-wafer interconnections. Lead-free soldershould be used for biocompatibility. Connection can also be achievedthrough IC processing techniques, which allow for plates and coils to beformed in electrical contact with one another. Finally laser welds, aspreviously discussed, can be used to achieve electrical/mechanicalinterconnects.

FIG. 16 illustrates a surface micromachined, capacitor coupled sensor600. The capacitor structure 602 comprises at least two plates 604, 606,at least one 604 of which is built directly atop a first wafer 608. Thisplate 604 will be referred to as the bottom plate. The region of thewafer 608 where the bottom plate 604 is built will be referred to as thedeflective region 610. If necessary, the thickness of the wafer 608 inthe region of the deflective region 610 can be reduced in thickness toenhance its deformability.

The other plate 606 is suspended above the bottom plate 604. The topplate 606 is mechanically anchored to the deflective region bypillar-like supporting elements 612 located at the periphery of thebottom plate 604. Bottom and top plates 604, 606 are electricallyinsulated and physically separated from one another by an air gap 614.The top electrode 606 mechanical design, material and dimensions arecarefully chosen so that the suspended part of the electrode does notstructurally deform under its own weight or creep over time.

A coil 616 of relevant geometry and inductance value is built orassembled using, as an example, any of the methods described herein. Itsterminals are electrically and mechanically connected to either one ofthe opposite plates 604, 606 of the capacitor 602. A capsule 618 orother form of hermetic surrounding is used to encapsulate both the coil616 and capacitor 602.

To achieve the desired pair of fixed and suspended plates 604, 606, thefabrication process of the disclosed embodiment employs a techniqueknown in the art as “sacrificial layer.” A sacrificial layer is astructural layer that remains buried throughout the fabrication processunder various layers of material until it can be removed, releasing thestructures and layers built on top of the sacrificial layer. Onceremoved, a void remains in place of the sacrificial layer. This voidforms the air gap that separates top from bottom plate(s).

A sacrificial layer must abide by at least two rules: (1) it must remainunaffected (no cracking, peeling, wrinkling, etc.) during the entirefabrication process until it is removed, and (2) selective and efficientremoval techniques must exist to remove it without adverse consequencesto any remaining structures.

Referring now to FIG. 17, the fabrication of the capacitor 602 startswith the creation of the bottom plate 604 on the wafer 808, usingphysical vapor deposition and photolithography. The backside of thewafer 608 is optionally thinned to enhance compliance in the deflectiveregion 610 of the wafer at the location of the bottom plate 604 so as tofacilitate deflection when a force or a pressure is applied.

The anchoring sites 612 are defined at the periphery of the bottom plate604. Anchoring sites 612 are small enough to represent only a fractionof the footprint of either bottom or top plate 604, 606. However, theyare big enough to insure reliable mechanical anchoring for the top plate606.

Referring now to FIG. 18, a layer 630 of material with desirablephysical and chemical traits is deposited onto the wafer 608 over thebottom plate 604 and the anchoring sites 612 to serve as a sacrificiallayer. The sacrificial material is, but is not limited to, a thin filmof photo-definable polymer (the first polymer layer). The thickness ofthe polymer is tuned by altering the conditions during deposition. Filmthicknesses ranging from fractions of micrometers to tens of micrometersare achieved routinely. To ensure that the layer 630 of photo-definablepolymer remains unaffected (no cracking, peeling, wrinkling, etc.)during the entire fabrication process until it is removed, proper curingand cross-linking precautionary steps must be taken.

With further reference to FIG. 18, using photolithography, windows 632are opened in the first polymer layer 630. The window geometry andin-plane location corresponds to those of the anchoring sites 612.Because the photo-definable polymer has a non-null thickness, eachopening (a.k.a. window) in the first polymer layer is surrounded bysidewalls 634 which height corresponds to the thickness of the firstpolymer layer.

A thin film metallic layer 640 is then deposited on top of thesacrificial layer 630, as depicted in FIG. 19. This layer comprises aseed layer, as it will provide a site upon which electroplated metalscan grow later on. The method of deposition should insure that themetallic film 640 evenly coats the upper surface of the sacrificiallayer 630 (the first polymer layer) as well as the sidewall 634 and thebottom areas of the windows 632 previously defined in the sacrificiallayer.

Referring now to FIG. 20, a second layer 650 of photo definable polymer(the second polymer layer) is deposited and patterned usingphotolithography. During this process, selected regions are removed fromthe surface of the substrate, defining new windows 652 (large openings)in the second polymer layer 650 without affecting any other previouslydeposited layer (especially the first polymer layer 630). The in-planegeometry of the new windows represents the in-plane geometry of the topelectrode 606 (FIG. 17). The geometry of the new windows extends toencompass the geometry and location of the anchor sites 612.

Regions where the photo definable polymer has been removed are subjectedto a method known as electroplating. In that fashion, metals like copperor gold can grow and adhere in the presence of the seed layer. Theelectroplating occurs at the same time at the anchoring sites, on thesidewalls, and on any other region exposed through windows opened in thesecond polymer layer. The resulting structure is a continuouselectroplated film 660 of the desired thickness. The thickness can rangefrom few micrometers to few tens of micrometers. Electroplated copper ispreferred for its ease of deposition and low cost.

Next, as shown in FIG. 21 the second polymer layer 650, the metal layer640, and the sacrificial layer 630 are removed using wet or dryselective removal techniques. The preferred removal technique for boththe second polymer layer 650 and the sacrificial layer 630 is wetdissolution in appropriate solvents such as acetone. At this point, bothbottom and top plates 604, 606 are formed. The top plate 606 issuspended above the bottom plate 604 and separated from it by an air gap614, which corresponds to the thickness of the first polymer layer.

As the fabrication of the sensor continues, the coil 616 is bunt orassembled using any of the methods described herein. Its terminals areelectrically and mechanically connected to either one of the oppositeplates 604, 606 of the capacitor 602. Finally, as shown in FIG. 16, thecapsule 618 or other form of hermetic surrounding is assembled onto thewafer 608 to encapsulate the coil 616 and capacitor 602.

A variation on the two-wafer design is shown in FIGS. 24-28. A sensor700 comprises a thick upper wafer 702 and a thinner lower wafer 704. Thethin lower wafer 704 comprises the pressure-sensitive deflective regionportion 706 of the sensor 700. A notch 708 is optionally formed in theupper wafer 702 to accommodate an anchor, such as a corkscrew, hook,barb, or other suitable stabilization means. The notch can be created onthe backside of the wafer directly if the cap is sufficiently thick toaccommodate the notch and a separation distance between the bottom ofthe notch and the coil body without causing any parasitic, deleteriouselectromagnetic or mechanical effects on the sensor function.Alternatively, the notch can be created by using wet or dry methods in aseparate wafer or plurality of wafers and then bonded to the backside ofthe sensor. The notch can have a variety of regular or irregulargeometries and can have rough or smooth sidewalls—any configurationachievable by conventional technologies that would impart some advantageor feature to assist in fixing the anchor mechanism to the sensor.

A capacitor 710 comprises a power plate 711 formed on the inner surfaceof the lower wafer 704 and an opposing pair of upper plates 712, 714formed on the lower surface of the upper wafer 702. A channel 716 isformed in the upper wafer 702 to receive an inductor coil 718. Theinductor coil 718 includes leads 720 that conductively connect theopposite ends of the coil to the upper plates 712, 714.

Manufacture of the sensor 700 will be explained with reference to FIGS.25-28. Referring first to FIG. 25, a dicing trench 730 is formed in thelower portion of the upper wafer 702 (shown inverted for themanufacturing process). The dicing trench 730 is a feature, whichcomprises a reduction in thickness of the wafer 702 along a line thatdefines the perimeter of the sensor 700. The dicing trench 730 isadvantageous where reduction of the amount of energy transferred to thesensor during dicing is needed, for example, to protect the sensor fromheat damage when dicing with a laser. When the wafer thickness isreduced, less energy is required to cut the sensor from the rest of thewafer, and thus less thermal energy is transferred to the criticalcomponents of the sensor.

As can also be seen in FIG. 25, the channel 716 is formed in the uppersurface of the upper wafer 702. The lower capacitor plates 712, 714 areformed on the upper surface of the upper wafer 702.

Referring now to FIG. 26, a recess 732 is formed in the upper surface ofthe lower wafer 704. The recess optionally includes troughs 734 forproviding clearance for the leads 720 of the inductor coil 718 (FIG.24). The lower capacitor plate 711 is formed in the base of the recess732 in the upper surface of the lower wafer 704.

Referring now to FIG. 27, the inductor coil 718 is introduced into theannular recess 716 of the upper wafer 702. The two leads 720 of theinductor coil 718 are connected to the upper capacitor plates 712, 714.

Referring to FIG. 28, the lower wafer 704 is now inverted and positionedatop the upper wafer 702. A laser is then used to cut and simultaneouslyheat bond the wafers 702, 704 at the lines 750 to complete fabricationof the sensor 700. Because of the presence of the dicing trenches 730,the laser need cut through only a thickness corresponding to the doublearrow 752. This shallow cut minimizes the amount of thermal energytransferred to the internal components of the sensor.

FIGS. 29-32 depict an embodiment of a sensor 800 manufactured from fourstacked wafers, 802, 804, 806, and 808. The bottom wafer 802 comprisesthe pressure-sensitive deflective region 810 and a pair of capacitorplates 812, 814 formed on its upper surface. The second wafer 804comprises a capacitor plate 816 formed on its lower surface and a pairof through-holes 818 for electrical connections. The third wafer 806comprises a cylindrical cavity 820 for accommodating an inductance coil822. Leads 824 of the inductance coil 822 extend through the holes 818in the second wafer 804 and connect to the capacitor plates 812, 814.The fourth wafer 808 fits atop the third wafer to provide a sealedstructure.

FIG. 30 illustrates a first step in the process for manufacturing thesensor 800. A recess 830 is formed in the upper surface of the bottomwafer. Then, as shown in FIG. 32, the plates 812, 814 are formed in thebase of the recess 830. Referring to FIG. 32, the plate 816 is formed onthe upper surface of the second wafer 804, and the through holes 818 areformed at the periphery of the plate 816. The second wafer is theninverted and stacked on top of the first wafer.

Thereafter, the coil 822 is positioned atop the second wafer, andelectrical connections are made through the holes 818 to the lowerplates 812, 814. After formation of the pressure sensitive, capacitorand inductor coil and connecting them together, hermetic encapsulationof the pressure sensitive cavity and inductor coil is performed. Thethird substrate wafer 806 is prepared with the deep recess 820,sufficient to contain the inductor coil 822. The recess 820 can beformed in a variety of ways, including laser rastering, glass machining,and ultrasonic machining. This third wafer 806 is bonded to the secondwafer 804 and subsequently, the sensors are cut out using a laser torelease the sensors from the wafer stack and form the hermetic seal inthe process of the cut.

The sensors described above can be adapted for use within an organ or alumen, depending upon what type of attachment or stabilizing means isemployed. FIGS. 33-36 illustrate a sensor 1001 suitable for use withinan organ such as the heart. The sensor 1001 has a generally cylindricalbody 1002 that hermetically houses the capacitor and inductor elementspreviously described. The sensor 1001 further has a pressure sensitivesurface 1003 (FIGS. 35 and 36) on one end of the cylindrical body 1002and a screw-type anchoring device 1004 extending upward from theopposite end of the body.

FIGS. 33-41 illustrate a first embodiment of a delivery device 1000(FIGS. 38, 40, and 41) for implanting a pressure sensor 1001 in a heartchamber. The sensor 1001 has a generally cylindrical body 1002 thathouses the capacitor and inductor elements previously described. Thesensor 1001 further has a pressure sensitive surface 1003 (FIGS. 35, 36,and 41) on one end of the cylindrical body 1002 and a screw-typeanchoring device 1004 extending upward from the opposite end of thebody. A retention mechanism 1005 of the delivery device 1000 comprises a“clamshell” housing 1006 wherein left and right housing halves 1008,1010 are resiliently deformable with respect to one another, much in themanner of a clothespin. The housing 1006 has a recess 1012 (FIGS. 35 and36) formed in its upper end, dimensioned to receive the sensor 1001therewithin. A reverse-threaded bore 1014 is formed in the lower end ofthe housing 1006, and a smooth counterbore 1016 is formed in the lowerend of the housing 1006 coaxially with the threaded bore 1014.

With further reference to the delivery device 1000, a screw 1018 has areverse-threaded shaft 1019 and a screw head 1020. The screw head 1020is mounted to the upper end of a dual-coil, flexible, torqueable shaft1022. As can be seen at 1024 of FIG. 37, a portion of the outer coil1026 is removed for purposes of illustration to show the inner coil1028, which is counterwound with respect to the outer coil 1026.

The reverse-threaded screw 1018 threadably engages the reverse-threadedbore 1014 in the lower end of the retention mechanism 1005. As the screwhead 1020 advances into the smooth counterbore 1016 in the base of thehousing 1006, the lower ends of the two housing halves 1008, 1010 arespread apart. This causes the upper ends of the housing halves 1008,1010 to dose together, thereby grasping the sensor 1001.

Referring now to FIGS. 38-41, delivery of the sensor 1001 of theinvention to a heart chamber may be accomplished as follows. Thephysician gains access into a vein that is suitable for access into theright ventricle using methods such as the Seldinger technique. Examplesof these access sites would be the right jugular, left subclavian, orright femoral veins. A guidewire is advanced into the right ventricle. Alarge vessel introducer with an adjustable hemostatic valve is insertedover the guidewire and advanced until its tip is positioned in the rightventricle.

The sensor 1001 is mounted to the delivery device 1000 with thelongitudinal axis of the device oriented normal to thepressure-sensitive surface of the sensor and with the anchor orstabilizer 1004 facing the distal end of the shaft 1022. The sensoranchor 1004 can be covered with a soluble, biocompatible material, or athin, retractable diaphragm cover (not shown). The purpose of suchcovering is to conceal the anchoring mechanism or stabilizer 1004 and toprotect the heart from inadvertent damage during sensor positioningprior to engaging the anchoring mechanism (which, in the case of thedisclosed sensor 1001 is configured to engage the tissue of the septum).A torqueable, kink-resistant, shaped guiding catheter (not shown) can beloaded over the shaft 1022 of the delivery device in order to provideadditional means for steering the sensor into position. Thecharacteristics of this guiding catheter are that the outer diameter issmall enough to fit within the introducer sheath, and the inner diameteris large enough to load over the shaft 1022 of the delivery device 1000.

Referring to FIG. 38, the shaft 1022 of the delivery device 1000 isrotated in a clockwise direction to screw the anchor 1004 of the sensorinto the tissue 1030 of the septum. When the anchor 1004 has been fullyinserted into the tissue 1030, as shown in FIG. 39, the sensor 1001tightens against the wall 1032 of the septum and creates a resistance.This resistance is sufficient to overcome the resistance between thereverse-threaded screw 1018 and the corresponding reverse-threaded bore1014 in the housing 1006 of the retention mechanism 1005. Consequently,continued rotation of the shaft 1022 of the delivery device 1000 in theclockwise direction will withdraw the screw 1018 from its bore 1014, asillustrated in FIG. 40. Once the screw head 1020 has cleared the smoothcounterbore 1016 in the lower end of the housing 1006 of the retentionmechanism, the lower ends of the two housing halves 1008, 1010 return totheir normal, closed configuration, thereby opening the upper ends ofthe two housing halves and releasing the sensor 1001, as depicted inFIG. 41. The delivery device 1000 is then withdrawn from the patient,leaving the sensor 1001 anchored to the wall 1032 of the septum with itspressure-sensing surface 1003 facing outward.

A feature of the disclosed embodiment is the use of a reverse-threadedscrew 1018 and corresponding bore 1014 so that rotating the shaft 1022in a normal “tightening” direction will first screw the sensor into thewall of the septum and then open the retention mechanism 1005 to releasethe sensor 1001, all without having to reverse direction of rotation ofthe shaft. To permit this arrangement, it is necessary that the screw1018 engage the retention mechanism 1005 with enough mechanical forcethat the initial rotation of the shaft 1022 will cause the sensor toscrew into the wall of the septum, rather than withdraw the screw 1018from the retention mechanism 1005. In addition, it is also necessarythat the screw be sufficiently loose with respect to the retentionmechanism that once the sensor has completely screwed into the wall ofthe septum, the torque resistance will overcome the engagement betweenthe screw and the retention mechanism rather than continue to rotate thesensor 1001. This feature can be accomplished, for example, bycontrolling the tolerances between the screw 1018 and the retentionmechanism 1005, and by controlling the resilient force exerted by thehousing 1006 against the head 1020 of the screw.

FIGS. 42 and 43 illustrate an alternate embodiment of a retentionmechanism 1055. The retention mechanism 1055 is mounted to a flexible,torqueable shaft 1022, just as in the previously disclosed embodiment.However, rather than the clamshell housing 1006, the retention mechanism1055 comprises a plurality of resilient wire fingers 1056 extendingupward from a base 1058. The fingers 1056 of the disclosed embodimentare comprised of nitinol, though any suitable resilient biocompatiblematerial can be used. Hooks 1060 at the upper ends of the wire fingers1056 wrap around the upper edges of the body 1002 of the sensor 1001. Inthe disclosed embodiment there are four such wire fingers 1056 spaced90.degree. apart around the circumference of the cylindrical sensor body1002, although a greater or lesser number of fingers 1056 can be used.Only two fingers 1056 are shown in the drawings for convenience ofillustration.

A spreader 1064 is disposed between the fingers 1056. The spreader 1064is attached to a pull-wire 1066, which extends through the longitudinalopening of the shaft 1022 and to a location outside of the patient. Whenthe physician desires to release the retention mechanism 1055 from thesensor 1001, he simply exerts a tension on the pull-wire 1066. Inresponse, the spreader moves downward and biases the fingers 1056 apart,releasing the sensor 1001 from the retention mechanism 1055. In thedisclosed embodiment the spreader 1064 is a circular disk or afrustocone, but it will be understood that any shape can be used whichbiases the fingers apart in response to tension applied to the pull-wire1066.

By changing the anchoring means, the same basic sensor 1001 can beadapted for use within a lumen such as an artery or arteriole in thepulmonary artery vasculature. FIGS. 44-46 illustrate a sensor 1100 ofthe type described above. The sensor 1100 has a wire loop 1102 extendingoutward from the sensor body 1104. As shown in FIG. 46, the wire loop1102 causes the sensor 1100 to lodge within a lumen 1106, with thesensor located centrally within the lumen and allowing blood flow allaround in the direction indicated by the arrow 1108.

A delivery apparatus for securing, delivering and deploying an implant1100 having an anchoring mechanism 1102 is shown in FIGS. 47-51. Thevarious components of the delivery apparatus are shown individually inFIGS. 47-50. As shown in FIG. 47, the delivery apparatus includes anelongated shaft 1152 having proximal and distal ends 1153, 1154respectively. The shaft 1152 has a main lumen 1155, which extends thelength of the shaft. A port 1156 places the main lumen 1155 incommunication with the ambient at an Intermediate location along theshaft 1152. A secondary lumen includes a proximal portion 1158 and adistal portion 1159. The proximal portion 1158 extends along a partiallength of the shaft 1152 and terminates in a port 1160 in the sidewallof the shaft. The distal portion 1159 originates in a port 1161 in thesidewall of the shaft and extends in a distal direction to an end.

A tether wire, 1163 shown in FIG. 48, is adapted to be slidablypositioned within the secondary lumen of the shaft 1152.

A core wire 1164, shown in FIG. 49, is configured to be received withinthe main lumen 1155 of the shaft 1152 and provides stiffness to thedelivery apparatus. The core wire 1164 has a decreasing diameter towardits distal end 1165, providing an increased flexibility in the distalend of the delivery apparatus. The core wire 1164 is fixed in the mainlumen 1155 of the shaft 1152 using adhesive, thermocompression, or anyother suitable fixation means.

Referring to FIG. 50, a conventional guide wire 1166 is dimensioned toextend beyond the distal end 1154 of the shaft 1152 and to be receivedwithin a distal portion of the main lumen 1155 of the shaft.

FIG. 51 shows the delivery apparatus with sensor 1100 mounted. The corewire 1164 is disposed within the main lumen 1155 of the shaft 1152. Thetether wire 1163 extends through the proximal portion 1158 of thesecondary lumen of the shaft 1152 and exits through the port 1160 in theshaft sidewall. The tether wire 1163 then is threaded through the body1104 of the sensor 1100 and passed into the port 1161 and hence into thedistal portion 1159 of the secondary lumen. The guidewire 1166 extendsalongside the proximal portion of the shaft 1152 and enters the mainlumen 1155 of the shaft 1152 at the port 1156. The guidewire 1166 thenpasses through the distal portion of the main lumen 1155 and exits thedistal end 1154 of the shaft 1152.

A vessel introducer is placed in an access site such as the rightinternal jugular vein, the subclavian artery, the right femoral vein, orany other suitable access site. The guidewire 1166 is inserted throughthe vessel introducer and guided to the target site using suitablemedical imaging technology. The delivery apparatus with sensor 1100mounted thereto is then threaded over the guidewire and inserted intothe vessel introducer.

After the delivery apparatus is in the vessel introducer, the apparatusis navigated over the guidewire to a deployment site in the pulmonaryartery. The implant 1100 is deployed by pulling the tether wire 1163proximally to disengage the implant from the shaft 1152. The deliveryapparatus and guidewire are then removed from the body.

The implant 1100 may then “float” through the narrowing pulmonary arteryvasculature until it reaches a location at which the vessel issufficiently narrow that the implant lodges within the vessel, as shownin FIG. 46. At that point the implant will be firmly anchored within thevasculature.

In alternate embodiments (not shown), the secondary lumen of the shaftcan comprise a single, uninterrupted lumen having two ports rather thantwo separate lumen portions. In addition, the secondary lumen can extendall the way through the distal end of the shaft, rather than terminatingat an end short of the distal end of the shaft.

Interrogation System

Embodiments are directed towards a system and method for communicatingwith a wireless sensor. Briefly described, the systems and methodsdetermines the resonant frequency of the sensor by adjusting the phaseand frequency of an energizing signal until the frequency of this signallocks to the resonant frequency of the sensor. The system energizes thesensor with a low duty cycle, gated burst of RF energy of apredetermined frequency or set of frequencies and predeterminedamplitude. This signal induces a current in the sensor that can be usedto track the resonant frequency of the sensor. The system receives thering down response of the sensor and determines the resonant frequencyof the sensor, which is used to calculate the measured physicalparameter. The system uses a pair of phase locked loops (“PLL”s) toadjust the phase and the frequency of the energizing signal to track theresonant frequency of the sensor.

Exemplary System

FIG. 52 illustrates an exemplary system for communicating with awireless sensor implanted within a body. The system includes a couplingloop 1200, a base unit 1202, a display device 1204 and an input device1206, such as a keyboard.

The coupling loop is formed from a band of copper. In one embodiment,the loop is eight inches in diameter. The coupling loop includesswitching and filtering circuitry that is enclosed within a shielded box1201. The loop charges the sensor and then couples signals from thesensor into the receiver. The antenna can be shielded to attenuatein-band noise and electromagnetic emissions.

Another possible embodiment for a coupling loop is shown in FIG. 58,which shows separate loops for energizing 1702 and for receiving 1704,although a single loop can be used for both functions. PIN diodeswitching inside the loop assembly is used to provide isolation betweenthe energizing phase and the receive phase by opening the RX path pindiodes during the energizing period, and opening the energizing path pindiodes during the coupling period. Multiple energizing loops can bestaggered tuned to achieve a wider bandwidth of matching between thetransmit coils and the transmit circuitry.

The base unit includes an RF amplifier, a receiver, and signalprocessing circuitry. Additional details of the circuitry are describedbelow in connection with FIG. 54.

The display 1204 and the input device 1206 are used in connection withthe user interface for the system. In the embodiment illustrated in FIG.52 the display device and the input device are connected to the baseunit. In this embodiment, the base unit also provides conventionalcomputing functions. In other embodiments, the base unit can beconnected to a conventional computer, such as a laptop, via acommunications link, such as an RS-232 link. If a separate computer isused, then the display device and the input devices associated with thecomputer can be used to provide the user interface. In one embodiment,LABVIEW software is used to provide the user interface, as well as toprovide graphics, store and organize data and perform calculations forcalibration and normalization. The user interface records and displayspatient data and guides the user through surgical and follow-upprocedures.

An optional printer 1208 is connected to the base unit and can be usedto print out patient data or other types of information. As will beapparent to those skilled in the art other configurations of the system,as well as additional or fewer components can be utilized with theinvention.

Patient and system information can be stored within a removable datastorage unit, such as a portable USB storage device, floppy disk, smartcard, or any other similar device. The patient information can betransferred to the physician's personal computer for analysis, review,or storage. An optional network connection can be provided to automatestorage or data transfer. Once the data is retrieved from the system, acustom or third party source can be employed to assist the physicianwith data analysis or storage.

FIG. 53 illustrates the system communicating with a sensor 1220implanted in a patient. The system is used in two environments: 1) theoperating room during implant and 2) the doctor's office duringfollow-up examinations. During implant the system is used to record atleast two measurements. The first measurement is taken duringintroduction of the sensor for calibration and the second measurement istaken after placement for functional verification. The measurements canbe taken by placing the coupling loop either on or adjacent to thepatient's back or the patient's stomach for a sensor that measuresproperties associated with an abdominal aneurysm. For other types ofmeasurements, the coupling loop may be placed in other locations. Forexample, to measure properties associated with the heart, the couplingloop can be placed on the patient's back or the patient's chest.

The system communicates with the implanted sensor to determine theresonant frequency of the sensor. As described in more detail in thepatent documents referenced in the Background section, a sensortypically includes an inductive-capacitive (“LC”) resonant circuithaving a variable capacitor. The distance between the plates of thevariable capacitor varies as the surrounding pressure varies. Thus, theresonant frequency of the circuit can be used to determine the pressure.

The system energizes the sensor with an RF burst. The energizing signalis a low duty cycle, gated burst of RF energy of a predeterminedfrequency or set of frequencies and a predetermined amplitude.Typically, the duty cycle of the energizing signal ranges from 0.1% to50%. In one embodiment, the system energizes the sensor with a 30-37 MHzfundamental signal at a pulse repetition rate of 100 kHz with a dutycycle of 20%. The energizing signal is coupled to the sensor via amagnetic loop. This signal induces a current in the sensor which hasmaximum amplitude at the resonant frequency of the sensor. During thistime, the sensor charges exponentially to a steady-state amplitude thatis proportional to the coupling efficiency, distance between the sensorand loop, and the RF power. FIG. 59 shows the charging response of atypical LC circuit to a burst of RF energy at its resonant frequency.The speed at which the sensor charges is directly related to the Q(quality factor) of the sensor. Therefore, the “on time” of the pulserepetition duty cycle is optimized for the Q of the sensor. The systemreceives the ring down response of the sensor via magnetic coupling anddetermines the resonant frequency of the sensor. FIG. 53A illustrates atypical energizing signal and FIGS. 53B, 53C and 53D illustrate typicalcoupled signals for various values of Q (quality factor) for the sensor.When the main unit is coupling energy at or near the resonant frequencyof the sensor, the amplitude of the sensor return is maximized, and thephase of the sensor return will be close to zero degrees with respect tothe energizing phase. The sensor return signal is processed viaphase-locked-loops to steer the frequency and phase of the nextenergizing pulse.

Operation of the Base Unit

FIG. 54 is a block diagram of the signal processing components within anexemplary base unit. The base unit determines the resonant frequency ofthe sensor by adjusting the energizing signal so that the frequency ofthe energizing signal matches the resonant frequency of the sensor. Inthe embodiment illustrated by FIG. 54, two separate processors 1302,1322 and two separate coupling loops 1340, 1342 are shown. In oneembodiment, processor 1302 is associated with the base unit andprocessor 1322 is associated with a computer connected to the base unit.In other embodiments, a single processor is used that provides the samefunctions as the two separate processors. In other embodiments a singleloop is used for both energizing and for coupling the sensor energy backto the receiver. As will be apparent to those skilled in the art, otherconfigurations of the base unit are possible that use differentcomponents.

The embodiment illustrated by FIG. 54 includes a pair of phase lockloops (“PLL”). One of the PLLs is used to adjust the phase of theenergizing signal and is referred to herein as the fast PLL. The otherPLL is used to adjust the frequency of the energizing signal and isreferred to herein as the slow PLL. The base unit provides two cycles:the calibration cycle and the measurement cycle. In one embodiment, thefirst cycle is a 10 microsecond energizing period for calibration of thesystem, which is referred to herein as the calibration cycle, and thesecond cycle is a 10 microsecond energizing/coupling period forenergizing the sensor and coupling a return signal from the sensor,which is referred to herein as the measurement cycle. During thecalibration cycle, the system generates a calibration signal for systemand environmental phase calibration and during the measurement cycle thesystem both sends and listens for a return signal, i.e. the sensor ringdown. Alternatively, as those skilled in the art will appreciate, thecalibration cycle and the measurement cycle can be implemented in thesame pulse repetition period.

The phase of the energizing signal is adjusted during the calibrationcycle by the fast PLL and the frequency of the energizing signal isadjusted during the measurement cycle by the slow PLL. The followingdescription of the operation of the PLLs is presented sequentially forsimplicity. However, as those skilled in the art will appreciate, thePLLs actually operate simultaneously.

Initially the frequency of the energizing signal is set to a defaultvalue determined by the calibration parameters of the sensor. Eachsensor is associated with a number of calibration parameters, such asfrequency, offset, and slope. An operator of the system enters thesensor calibration parameters into the system via the user interface andthe system determines an initial frequency for the energizing signalbased on the particular sensor. Alternatively, the sensor calibrationinformation could be stored on portable storage devices, bar codes, orincorporated within a signal returned from the sensor. The initial phaseof the energizing signal is arbitrary.

The initial frequency and the initial phase are communicated from theprocessor 1302 to the DDSs (direct digital synthesizers) 1304, 1306. Theoutput of DDS1 1304 is set to the initial frequency and initial phaseand the output of DDS2 1306 (also referred to as local oscillator 1) isset to the initial frequency plus the frequency of the local oscillator2. The phase of DDS2 is a fixed constant. In one embodiment, thefrequency of local oscillator 2 is 4.725 MHz. The output of DDS1 isgated by the field programmable gate array (FPGA) 1308 to create apulsed transmit signal having a pulse repetition frequency (“PRF”). TheFPGA provides precise gating so that the base unit can sample thereceive signal during specific intervals relative to the beginning orend of the calibration cycle.

During the calibration cycle, the calibration signal which enters thereceiver 1310 is processed through the receive section 1311 and the IFsection 1312, and is sampled. In one embodiment, the calibration signalis the portion of the energizing signal that leaks into the receiver(referred to herein as the energizing leakage signal). The signal issampled during the on time of the energizing signal by a sample and holdcircuit 1314 to determine the phase difference between the signal andlocal oscillator 2. In the embodiment where the calibration signal isthe portion of the energizing signal that leaks into the receiver, thesignal is sampled approximately 100 ns after the beginning of theenergizing signal pulse. Since the energizing signal is several ordersof magnitude greater than the coupled signal, it is assumed that thephase information associated with the leaked signal is due to theenergizing signal and the phase delay is due to the circuit elements inthe coupling loop, circuit elements in the receiver, and environmentalconditions, such as proximity of reflecting objects.

The phase difference is sent to a loop filter 1316. The loop filter isset for the dynamic response of the fast PLL. In one embodiment, the PLLbandwidth is 1000 Hz and the damping ratio is 0.7. A DC offset is addedto allow for positive and negative changes. The processor 1302 reads itsanalog to digital converter (ND) port to receive the phase differenceinformation and adjusts the phase sent to direct digital synthesizer 1(DDS1) to drive the phase difference to zero. This process is repeatedalternatively until the phase difference is zero or another referencephase.

The phase adjustment made during the energizing period acts to zero thephase of the energizing signal with respect to local oscillator 2.Changes in the environment of the antenna or the receive chainimpedance, as well as the phase delay within the circuitry prior tosampling affect the phase difference reading and are accommodated by thephase adjustment.

During the measurement cycle, the energizing signal may be blocked fromthe receiver during the on time of the energizing signal. During the offtime of the energizing signal, the receiver is unblocked and the coupledsignal from the sensor (referred to herein as the coupled signal or thesensor signal) is received. The coupled signal is amplified and filteredthrough the receive section 1311. The signal is down converted andadditional amplification and filtering takes place in the IF section1312. In one embodiment, the signal is down converted to 4.725 MHz.After being processed through the IF section, the signal is mixed withlocal oscillator 2 and sampled by sample and hold circuits 1315 todetermine the phase difference between the coupled signal and theenergizing signal. In one embodiment, the sampling occurs approximately30 ns after the energizing signal is turned off.

In other embodiments, group delay or signal amplitude is used todetermine the resonant frequency of the sensor. The phase curve of asecond order system passes through zero at the resonant frequency. Sincethe group delay i.e. derivative of the phase curve reaches a maximum atthe resonant frequency, the group delay can be used to determine theresonant frequency. Alternatively, the amplitude of the sensor signalcan be used to determine the resonant frequency. The sensor acts like abandpass filter so that the sensor signal reaches a maximum at theresonant frequency.

The sampled signal is accumulated within a loop filter 1320. The loopfilter is set for the dynamic response of the slow PLL to aid in theacquisition of a lock by the slow PLL. The PLLs are implemented withop-amp low pass filters that feed ND inputs on microcontrollers, 1302and 1322, which in turn talk to the DDSs, 1304 and 1306, which providethe energizing signal and local oscillator 1. The microcontroller thatcontrols the energizing DDS 1304 also handles communication with thedisplay. The response of the slow PLL depends upon whether the loop islocked or not. If the loop is unlocked, then the bandwidth is increasedso that the loop will lock quickly. In one embodiment, the slow PLL hasa damping ratio of 0.7 and a bandwidth of 120 Hz when locked (theNyquist frequency of the blood pressure waveform), which isapproximately ten times slower than the fast PLL.

A DC offset is also added to the signal to allow both a positive and anegative swing. The output of the loop filter is input to an ND input ofprocessor 1322. The processor determines a new frequency and sends thenew frequency to the DSSs. The processor offsets the current frequencyvalue of the energizing signal by an amount that is proportional to theamount needed to drive the output of the slow PLL loop filter to apreset value. In one embodiment the preset value is 2.5V and zero inphase. The proportional amount is determined by the PLL's overalltransfer function.

The frequency of the energizing signal is deemed to match the resonantfrequency of the sensor when the slow PLL is locked. Once the resonantfrequency is determined, the physical parameter, such as pressure, iscalculated using the calibration parameters associated with the sensor,which results in a difference frequency that is proportional to themeasured pressure.

The operation of the slow PLL is qualified based on signal strength. Thebase unit includes signal strength detection circuitry. If the receivedsignal does not meet a predetermined signal strength threshold, then theslow PLL is not allowed to lock and the bandwidth and search window forthe PLL are expanded. Once the received signal meets the predeterminedsignal strength threshold, then the bandwidth and search window of theslow PLL is narrowed and the PLL can lock. In the preferred embodiment,phase detection and signal strength determination are provided via the“I” (in phase) and “Q” (quadrature) channels of a quadrature mixercircuit. The “I” channel is lowpass filtered and sampled to providesignal strength information to the processing circuitry. The “Q” channelis lowpass filtered and sampled to provide phase error information tothe slow PLL.

Avoiding False Locks

The system provides unique solutions to the false lock problem. A falselock occurs if the system locks on a frequency that does not correspondto the resonant frequency of the sensor. There are several types offalse locks. The first type of false lock arises due to the pulsednature of the system. Since the energizing signal is a pulsed signal, itincludes groups of frequencies. The frequency that corresponds to afalse lock is influenced by the pulse repetition frequency, the Q of thesensor, and the duty cycle of the RF burst. For example, a constantpulse repetition frequency adds spectral components to the return signalat harmonic intervals around the resonant frequency of the sensor, whichcan cause a false lock. In one embodiment, false locks occur atapproximately 600 kHz above and below the resonant frequency of thesensor. To determine a false lock, the characteristics of the signal areexamined. For example, pulse repetition frequency dithering and/orobserving the slope of the baseband signal are two possible ways ofdetermine a false lock. In one embodiment where the system locks on asideband frequency, the signal characteristics correspond to a heartbeator a blood pressure waveform.

The second type of false lock arises due to a reflection or resonance ofanother object in the vicinity of the system. This type of false lockcan be difficult to discern because it generally does not correspond toa heartbeat or blood pressure waveform. The lack of frequency modulationcan be used to discriminate against this type of false lock. Changingthe orientation of the magnetic loop also affects this type of falselock because the reflected false lock is sensitive to the angle ofincidence. The third type of false lock arises due to switchingtransients caused by switching the PIN diodes and analog switches in theRF path. These transients cause damped resonances in the filters in thereceive chain, which can appear similar to the sensor signal. Typically,these types of false locks do not correspond to a heartbeat or bloodpressure waveform because they are constant frequency. These types offalse locks are also insensitive to orientation of the magnetic loop.

To avoid the first type of false lock, the embodiments herein determinethe slope of the baseband signal (the phase difference signal at point1330). In one embodiment, if the slope is positive, then the lock isdeemed a true lock. However, if the slope is negative, then the lock isdeemed a false lock. In another embodiment, a negative slope is deemed atrue lock and a positive slope is deemed a false lock. The slope isdetermined by looking at points before and after the phase differencesignal goes to zero. The slope can be determined in a number ofdifferent ways, including but not limited to, using an analogdifferentiator or multiple sampling. FIGS. 55A and 55B illustrate a truelock and a false lock respectively, when a positive slope indicates atrue lock. In one embodiment, if a false lock is detected, then thesignal strength is suppressed so that the signal strength appears to theprocessor 1322 to be below the threshold and the system continues tosearch for the center frequency. In other embodiments, any non-zeroslope can be interpreted as a false lock resulting in zero signalstrength.

The system can also use frequency dithering to avoid the first type offalse lock. Since the spectral components associated with a constantpulse repetition frequency can cause a false lock, dithering the pulserepetition frequency helps avoid a false lock. By dithering the pulserepetition frequency, the spectral energy at the potential false lockfrequencies is reduced over the averaged sampling interval. As shown inFIG. 56, the energizing signal includes an on time t1 and an off timet2. The system can vary the on time or the off time to vary the PRF(PRF=1/(t1+t2)). FIG. 56 illustrates different on times (t1, t1′) anddifferent off times (t2, t2′). By varying the PRF, the sidebands moveback and forth and the average of the sidebands is reduced. Thus, thesystem locks on the center frequency rather than the sidebands. The PRFcan be varied between predetermined sequences of PRFs or can be variedrandomly.

Reducing Switching Transients

The coupling loop switches between an energizing mode and a couplingmode. This switching creates transient signals, which can cause thethird type of false lock. Phase dithering is one method used to reducethe switching transients. As shown in FIG. 57, the system receives aswitching transient 1603 between the end of the energizing signal 1602and the beginning of the coupled signal 1604. To minimize the transient,the phase of the energizing signal may be randomly changed. However,changing the phase of the energizing signal requires that the systemredefine zero phase for the system. To redefine zero phase for thesystem, the phase of DDS2 is changed to match the change in phase of theenergizing signal. Thus, the phase of the energizing signal 1602′ andthe coupled signal 1604′ are changed, but the phase of the transientsignal 1603′ is not. As the system changes phase, the average of thetransient signal is reduced.

Changing the resonant frequency of the antenna as it is switched fromenergizing mode to coupling mode also helps to eliminate the switchingtransients. Eliminating the switching transients is especially importantin the present invention because of the characteristics of the coupledsignal. The coupled signal appears very quickly after the on period ofthe energizing signal and dissipates very quickly. In one embodiment,the invention operates in a low power environment with a passive sensorso that the magnitude of the coupled signal is small. However, theinvention is not limited to working with a passive sensor.

The coupling loop is tuned to a resonant frequency that is based uponthe sensor parameters. Changing the capacitors or capacitor network thatis connected to the coupling loop changes the resonant frequency of theantenna. The resonant frequency typically is changed from approximately1/10% to 2% between energizing mode and coupled mode. In someembodiments, the coupling loop is untuned.

Additional alternative embodiments will be apparent to those skilled inthe art to which the present invention pertains without departing fromits spirit and scope. For example, the system can operate with differenttypes of sensors, such as non-linear sensors that transmit informationat frequencies other than the transmit frequency or sensors that usebackscatter modulations. Accordingly, the scope of the present inventionis described by the appended claims and is supported by the foregoingdescription.

Finally, it will be understood that the preferred embodiment has beendisclosed by way of example, and that other modifications may occur tothose skilled in the art without departing from the scope and spirit ofthe appended claims.

What is claimed is:
 1. A method for determining a pressure associatedwith a lumen of a body, comprising: positioning a wireless sensor in thelumen of the body, the sensor comprising an LC resonant circuit having aresonant frequency configured to vary in response to changes in pressurein the lumen; storing one or more sensor calibration parameters at anexternal base unit; generating an energizing signal from the externalbase unit; transmitting the energizing signal; receiving a ring downresponse from the wireless sensor; determining the resonant frequency ofthe LC resonant circuit from the ring down response; and calculating thepressure in the lumen from the resonant frequency of the LC resonantcircuit utilizing the one or more sensor calibration parametersassociated with the LC resonant circuit.
 2. The method of claim 1,further comprising obtaining the one or more sensor calibrationparameters from a bar code associated with the wireless sensor.
 3. Themethod of claim 1, further comprising obtaining the one or more sensorcalibration parameters from a portable storage device.
 4. The method ofclaim 1, further comprising obtaining the one or more sensor calibrationparameters incorporated within a signal returned from the wirelesssensor.
 5. The method of claim 1 further comprising distributing atleast one of i) the one or more sensor calibration parameters or ii)patient information, to multiple devices using a network or one or moreportable storage devices.
 6. The method of claim 1, wherein the one ormore sensor calibration parameters includes at least one of a frequency,offset, or slope associated with the wireless sensor.
 7. The method ofclaim 1, further comprising taking a pressure measurement during atleast one of i) when introducing the wireless sensor or ii) afterplacement of the wireless sensor.
 8. The method of claim 7, furthercomprising utilizing the pressure measurement for calibration of thewireless sensor.
 9. The method of claim 8, utilizing the one or morepressure calibration parameters for the calibration of the wirelesssensor.
 10. The method of claim 1, wherein the generating comprises:generating the energizing signal during a measurement cycle and gatingthe energizing signal such that the duty cycle has an on-time that isset based on a speed at which the LC resonant circuit charges.
 11. Themethod of claim 1, wherein the lumen of the body is an artery.
 12. Themethod of claim 1, wherein the ring down response is received during atleast one of a calibration cycle or the measurement cycle.
 13. Themethod of claim 1, further comprising placing one or more coupling loopsin relative alignment with the implanted sensor, the one or morecoupling loops connected to the external base unit, the energizingsignal generated at the one or more coupling loops, the energizingsignal comprises low duty cycle radio frequency (RF) energy having a setof two or more frequencies, each of the energizing signals having acenter frequency staggered with respect to one another.
 14. A system,comprising: an implantable wireless sensor for determining a pressure ofa lumen of a body, the wireless sensor comprising an LC resonant circuithaving a resonant frequency configured to vary in response to changes inpressure in the lumen; an external base unit connected to the couplingloop, the external base unit configured to: store one or more sensorcalibration parameters into the base unit; generate an energizingsignal; transmit the energizing signal; receive a ring down responsefrom the wireless sensor; determining the resonant frequency of the LCresonant circuit from the ring down response; and calculate the pressurein the lumen from the resonant frequency of the LC resonant circuitutilizing the one or more sensor calibration parameters associated withthe LC resonant circuit.
 15. The system of claim 14, wherein theexternal base unit is further configured to obtain the one or moresensor calibration parameters from a bar code associated with thewireless sensor
 16. The system of claim 14, wherein the external baseunit is further configured to obtain the one or more sensor calibrationparameters from a portable storage device.
 17. The system of claim 14,wherein the external base unit is further configured to obtain the oneor more sensor calibration parameters from a signal returned from thewireless sensor.
 18. The system of claim 14, wherein the external baseunit is further configured to distribute at least one of i) the one ormore sensor calibration parameters or ii) patient information, tomultiple devices using a network or one or more portable storagedevices.
 19. The system of claim 14, wherein the one or more sensorcalibration parameters includes at least one of a frequency, offset, orslope associated with the wireless sensor.
 20. The system of claim 14,wherein the external base unit is further configured to obtain apressure measurement during at least one of i) when introducing thewireless sensor or ii) after placement of the wireless sensor.
 21. Thesystem of claim 20, wherein the external base unit is further configuredto utilize the pressure measurement for calibration of the wirelesssensor.
 22. The system of claim 21, wherein the external base unit isfurther configured to utilize the one or more pressure calibrationparameters for the calibration of the wireless sensor.
 23. The system ofclaim 14, wherein the energizing signal comprises low duty cycle radiofrequency (RF) energy having a set of two or more frequencies, each ofthe energizing signals having a center frequency staggered with respectto one another.
 24. The system of claim 14, further comprising acoupling loop connected to the external base unit, the coupling loopconfigured to be placed in relative alignment with the implanted sensorand to generate the energizing signal.